Cancer is a leading cause of death in the United States. Each year, more than a million U.S. residents are diagnosed with cancer; however, recent statistics indicate that for the first time in five decades, cancer mortality is declining. Advances in cancer diagnostic techniques involving imaging technology is one of the key factors contributing to this decrease. Currently, the standard procedure for cancer diagnosis requires a biopsy of a suspect site, followed by tissue histology. Unfortunately, biopsy is an invasive procedure and often results in unacceptably high rates of false negative diagnoses because of random sampling errors, particularly when the area of interest is small (as in cases of early cancers). Thus, it would be preferable to use a different approach for cancer screening and early cancer detection that provides more accurate results and is less invasive.
More than 70% of all cancers originate in the epithelial lining of internal organs. Some of the more common examples include cancers of the esophagus, colon, bladder, and lung that can develop over a period of several years and are characterized by changes in tissue and cellular morphology before invasion and metastasis occur. While x-rays, positron emission tomography (PET), magnetic resonance imaging (MRI), ultrasound, and surface tissue endoscopy have all played significant roles in the detection of macroscopic abnormalities (e.g., large tumors and strictures), physicians are still challenged by their limited ability to detect and examine microscopic changes in early-stage neoplasia in vivo, using current clinical imaging technologies, which often have insufficient image resolution to provide the information required.
Scanning confocal microscopy, OCT, and multiphoton microscopy (MPM) are three non-invasive optical technologies that are capable of imaging tissue microstructures at or near cellular resolution (˜0.5-10 μm) in vivo. These technologies have the potential for performing “optical biopsy” at resolutions near those of conventional histological techniques, but without the need for tissue removal. All three techniques require mechanisms that deliver, focus, scan, and collect a single mode optical beam. Conventional microscopes equipped with galvanometer or rotating polygon mirrors can perform this scanning task when imaging biological samples or tissues that are easily accessed externally, outside a patient's body. Yet, such devices are typically too bulky and are thus often the limiting factor in imaging probe miniaturization. Imaging internal organs requires extreme miniaturization of the scanning apparatus. Although it is possible to deliver an optical beam to internal organs using a single mode optical fiber, the integration of beam scanning, focusing, and collection using an endoscope only a few millimeters in diameter is a major engineering challenge that has not successfully been solved in the prior art.
A further challenge is the need for a focus-tracking mechanism that can maintain a high transverse resolution at varying depths, in particular when the focused spot sizes are small. For example, conventional confocal microscopy is a well-established technique that can image tissue specimens and living tissues at cellular resolution, and most in vivo human imaging using confocal microscopy focuses on tissues that are easily accessed externally, such as the eyes and skin. Although attempts have been made to integrate fiber-optic imaging bundles with scanning confocal microscopy for imaging internal organs, wherein the fiber-optic bundle relays tissue images from internal organs to a conventional scanning confocal microscope outside the human body, the resulting resolution is sub-optimal and generally unsatisfactory for most purposes. The lower resolution is primarily due to cross-talk between fibers and limited fiber packing density.
Recently, micro-electrical-mechanical-system (MEMS) scanners have undergone intensive investigation, and it appears that it may be possible to use MEMS scanners to perform beam scanning endoscopically. Yet, a MEMS-scanner-based endoscope is still relatively large (e.g., ˜5-8 mm in diameter) because of the required supporting substrate, electrodes, and packaging. In addition, MEMS scanners may also introduce wave front deformation to the imaging beam, since MEMS mirrors are thin and tend to warp during scanning.
Perhaps a more promising approach in endoscopic beam scanning is to scan an optical fiber tip to image tissue at a desired internal location within a patient's body. An optical fiber can be mounted on a metal base plate (e.g., a tuning fork) and actuated by electromagnetic oscillation. An imaging device using this scanning scheme with a diameter of ˜3-6 mm has been demonstrated. Further size reduction is difficult, as a result of limitations imposed by the size of the electromagnetic actuator. A similar approach has been reported in which an optical fiber attached to an electric coil is actuated by a stationary magnet when an AC current is applied to the coil. This optical fiber scanner has a smaller diameter, ˜3 mm, and can achieve a 2-mm transverse scan. Yet, the scanning speed is severely limited to a few transverse scans per second.
Providing suitable miniature imaging optics is another important consideration for confocal endoscopy. Elegant miniature optics using a graded index (GRIN) lens and a compound sol-gel lens have been reported in the prior art. The magnification achieved by such a device is ˜4×-8×, corresponding to a focal spot size of about 20-40 μm (i.e., a fiber mode field diameter of 5 μm, multiplied by the magnification factor). Although GRIN lenses can be readily implemented in the scanning fiber endoscope, it is well known that GRIN lenses can cause chromatic aberrations. Yet, this problem can likely be resolved by developing miniature optics with a lower magnification as well as minimal optical aberration (spherical and chromatic).
OCT is an emerging non-invasive technology that can perform cross-sectional imaging of tissue microstructures in vivo and in real-time. OCT is analogous to ultrasound in imaging applications, except that it uses low-coherence light rather than acoustic waves to image tissues. The echo delay time or the depth of light backscattered from the tissue is measured using a technique called low coherence interferometry. The heterodyne detection gives OCT extremely high detection sensitivity in excess of 100 dB, corresponding to the detection of backscattered optical signals of 1 part in 1010. FIG. 1 (Prior Art) schematically illustrates a conventional OCT system. This system includes a Michelson interferometer that uses a low coherence light source 20. The light source is coupled to an OCT probe 24 in the sample arm and to a reference arm 28 through an optic fiber coupler or beam splitter 22. The sample arm delivers an optical beam from the light source to tissue 26 and collects the backscattered light. The reference arm performs depth scanning by using a translating retro-reflective mirror or a phase-controlled scanning delay line (not separately shown). The backscattered intensity versus depth forms an axial scan. Two- or three-dimensional data sets formed by multiple adjacent axial scans are obtained by scanning the OCT beam along the transverse direction after each axial scan. A photodetector 30 produces a corresponding analog signal comprising the data set. The analog signal is processed by detection electronics module 32, which produces corresponding digital data. The resulting data set can be displayed using a computer 38, as a false-color or gray-scale map, to form a cross-sectional OCT image.
Unlike confocal microscopy, the transverse and axial resolutions of OCT are determined independently. The axial resolution, Δz, is given by the coherence length of the light source and is inversely proportional to the source spectrum bandwidth Δλ, i.e., Δz=(2 ln 2/π)(λ02/Δλ), where λ0 is the source center wavelength. The transverse resolution, Δx, is determined by the transverse focused spot size, in a manner similar to that in conventional microscopy, i.e., Δx=(2λ/π)/N.A., where N.A.=d/2f, d is the beam spot size on the objective lens, and f is the focal length of the objective.
It is well known that an increase in the transverse resolution reduces the depth of focus quadratically, i.e., b=(πΔx2)/2λ, where b is the depth of focus (or the confocal parameter). For example, the depth of focus decreases from ˜200 μm to ˜50 μm when the transverse resolution increases from 10 μm to 5 μm. Conventional OCT has a low transverse resolution between 20 μm and 40 μm. Thus, focus tracking is not necessary for low resolution OCT. However, low transverse resolution degrades image contrast. Even with coherence gating along the axial direction, photons that are backscattered within the focal spot size by different scatterers (e.g., by cells or cell organelles) will likely be simultaneously detected and averaged, causing loss of contrast. Therefore, a high transverse resolution is needed. When high N.A. optics are utilized to achieve a high transverse resolution, focus tracking is clearly required. As discussed above, conventional OCT imaging acquires one axial scan followed by other axial scans, each at a different transverse location. A 2-3 mm axial scan generally takes less than 2 ms during real-time imaging, requiring focus tracking at a velocity of ˜4-6 meters per second, which is extremely difficult to achieve in a compact scanning device. FIG. 2 illustrates the rapid depth scanning of tissue 42 by an incident beam 40, and the relatively slow transverse scans that are used. In this conventional technique for OCT scanning, focus tracking means that the focus point is rapidly tracked at each different transverse location before moving to the next transverse location, which is very challenging to achieve.
The core of a single mode optical fiber in an OCT imaging system will function optically much as a pinhole does in a confocal microscope. Thus, a fiber-optic OCT system can also be employed as a confocal microscope, with the added benefit that OCT provides superb axial resolution by using coherence gating. Reportedly, a confocal microscope equipped with low coherence gating (resulting in a device known as an optical coherence microscope) improves the imaging depth by a factor of more than 2, compared to conventional confocal microscopy. A unified imaging modality is expected to have an enhanced resolution (both transverse and axial) and should enable imaging tissue microstructures at or near cellular levels.
Clearly, what is needed is an approach that enables forward-directed OCT scanning or confocal imaging to be carried out without the need for focus tracking at high velocities that are difficult to achieve. An OCT scanning system and technique would be desirable that can be implemented using a scanner sufficiently compact to be readily inserted within a patient's body with minimal invasive consequences. The scanner should produce high resolution scans and provide detailed information that can be evaluated by medical personnel to determine the condition of the tissue being imaged. Prior art advances in OCT and confocal imaging have not yet achieved this goal.
There is another important related imaging paradigm. Multiphoton microscopy (MPM) has become a powerful tool for detailing subcellular structures and events. This technology relies on two or more long wavelength photons arriving “simultaneously” at a fluorophore, where the energies add and induce an electronic transition that is normally excited by a single short wavelength photon. Among the many advantages over single photon excitation, the nonlinear excitation process of MPM is restricted to a submicron-size volume at the focus of the light beam, providing a superb resolution. Unlike OCT and confocal microscopy, MPM is sensitive to biochemical information, including cellular NAD(P)H, flavin, retinal condition, etc. Recent in vivo animal model studies have demonstrated that MPM is an enabling technology for assessing tumor pathophysiology and differentiating metastatic from non-metastatic tumors. A variation of MPM may become useful for imaging anisotropic molecules and biological structures without the requirement of fluorescence; this variation employs the mechanism of harmonic generation such as second harmonic generation (SHG) or higher harmonic generation. For example, the SHG signal from collagen is typically at one-half the two-photon excitation wavelength in the near infrared. Studies have also shown the feasibility of MPM for in vivo imaging of human skin with cellular resolution.
Recently, active research has been devoted to endoscopic MPM. One major technical difficulty is the temporal broadening of femtosecond pulses through optical fibers due to the group velocity dispersion (GVD) and self-phase modulation (SPM), which results in a power-law decrease of multiphoton excitation efficiency. Preliminary studies have suggested that this problem could be potentially overcome using large core multimode fibers, novel microstructured or photonic bandgap optical fibers. It would thus also be desirable to develop endoscopic applications of MPM using an optical fiber scanner.